Charge perturbation detection system for dna and other molecules

ABSTRACT

Methods and apparatus for direct detection of chemical reactions are provided. Electric charge perturbations of the local environment during enzyme-catalyzed reactions are sensed by an electrode system with an immobilized target molecule. The charge perturbation caused by the polymerase reaction can uniquely identify a DNA sequence. The polymerization process generates local perturbations of charge in the solution near the electrode surface and induces a charge in a polarazible gold electrode. This event is detected as a transient current by a voltage clamp amplifier. Detection of single nucleotides in a sequence can be determined by dispensing individual dNTPs to the electrode solution and detecting the charge perturbations. Alternatively, multiple bases can be determined at the same time using a mix of all dNTPs with subsequent analysis of the resulting signal. This technique may be adapted to other reaction determinations, such as enzymatic reactions, other electrode configurations, and other amplifying circuits.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.13/434,627 filed on Mar. 29, 2012, which is a divisional of U.S. patentapplication Ser. No. 13/170,607 filed on Jun. 28, 2011, now U.S. Pat.No. 8,313,907, which is a continuation of U.S. patent application Ser.No. 12/821,809 filed on Jun. 23, 2010, now U.S. Pat. No. 8,012,756,which is a continuation of U.S. patent application Ser. No. 11/271,678filed on Nov. 10, 2005, now U.S. Pat. No. 7,785,785, which claimspriority from U.S. Provisional Patent Application No. 60/627,192 filedon Nov. 12, 2004, all of which are hereby incorporated by reference intheir entirety.

STATEMENT OF GOVERNMENTAL SUPPORT

This invention was made with Government support under contracts AI059499and HG000205 awarded by the National Institutes of Health. TheGovernment has certain rights in this invention.

REFERENCE TO SEQUENCE LISTING OR CD ROM

Applicants submit herewith a sequence listing in an ASCII text file(3815_(—)10_(—)5_seq_list.txt), as provided in EFS Legal FrameworkNotice 20 May 2010, part I-I-1. The file was created Oct. 18, 2012 andcontains 834 bytes. Applicants incorporate the contents of the sequencelisting by reference in its entirety.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to the field of electrochemical moleculardetection, such as the detection of nucleic acid polymerization, bydetection of a charged particle (e.g. proton) by means of a sensitiveelectrical circuit. The present invention has particular application tonucleic acid sequence detection and enzyme substrate modificationdetection.

2. Related Art

Rapid, sequence-specific DNA detection is essential for applications inmedical diagnostics and genetic screening. Electrochemical biosensorsthat use immobilized nucleic acids are especially promising in theseapplications because of their potential for miniaturization andautomation. Current DNA detection methods based on hybridization rely onvarious optical, electrochemical or mass readouts. (Refs. 1, 2) However,direct, label-free electrochemical detection methods are not available.

A variety of electrochemical methods have been described, all of whichdetect direct electronic signals using various electrochemical reactionsduring DNA hybridization at the electrode surface. (Refs. 1, 2, 6) Incontrast, current detection and sequencing-by-synthesis techniquerequires the use of several enzymatic and photochemical steps. (Refs.7,8) Thus a direct electrochemical detection method for this techniquewould greatly simplify the detection process and accelerate itsimplementation for rapid DNA sequencing and diagnostics. Described belowis a label-free electrochemical detection method, Charge PerturbationDetection (CPD), applied to sequencing-by-synthesis.

DISCUSSION OF RELATED PUBLICATIONS AND PATENTS

Drummond, T. G., M. G. Hill, J. K. Barton, “Electrochemical DNAsensors,” Nat. Biotechnol. October; 21(10):1192

9 (2003) report direct electrochemical techniques based on detection ofelectronic signals of electrochemical reactions of DNA or reportermolecules or enzymes recruited to the electrode surface by specific DNAprobe-target interactions.

Ronaghi, M. “Pyrosequencing sheds light on DNA sequencing,” Genome Res.Ian; 11 (1):3-11 (2001) discloses pyrosequencing detection methods usinga bioluminometric detection in a three step reporter technique. Theultimate goal of these efforts is to discriminate individual nucleotidesin a DNA molecule. Almost all these techniques use more than one-stepelectrochemical reaction to produce electronic signal. The CPD (Chargeperturbation detection) method and device described below in contrastuses only one-step polymerase-catalyzed reaction to generate anelectronic signal detecting a DNA nucleotide or sequence.

Pourmand et al. US 2002/0155476, published Oct. 24, 2002, describes adevice for detecting a transient electrical signal in a sample. Thisdevice relies on a change in the potential difference between twoelectrodes when ions are added to a medium contacting the electrodes. Adifferential amplifier subtracts voltages from the two electrodes toproduce the signal. The signal is generated in response to an electricfield generated by the migration of ions towards the binding sites onone electrode, as shown in FIG. 2.3 of the publication.

BRIEF SUMMARY OF THE INVENTION

The invention comprises a method and device for direct electrochemicaldetection of enzymatically catalyzed DNA synthesis by induced surfacecharge perturbation. Incorporation of a complementary nucleotidetriphosphate (NTP) such as deoxynucleotide triphosphate (dNTP) into aself-primed single-stranded DNA attached to the surface of a highoverpotential metal (e.g. gold) electrode evokes an electrode surfacecharge perturbation. This event can be detected as a transient currentby a feed back (e.g. voltage-clamp) amplifier.

Based on current understanding of polarizable interfaces (Refs 3, 4) andwithout wishing to be bound by scientific theory, it is thought that theelectrode detects the proton removal from the 3′-hydroxyl group of theDNA molecule during phosphodiester bond formation. (Ref 5).

Thus the present invention provides a device for detecting a chemicalreaction of a selected reactant with a target molecule, wherein thereaction produces a charge perturbation on the target molecule. It neednot be limited to DNA or other nucleic acid polymerization reactions.The chemical reaction may further provide structural information, suchas a DNA or RNA sequence. It may provide information about the state ofphosphorylation of a substrate, etc. The device comprises a containerfor containing reaction medium having therein reactants and targetmolecules. The container may be a well or a channel; the targetmolecules (e.g. DNA) may be chemically linked to the electrode, orsimply in suspension near the electrode. Since the detection zone in thepreferred embodiment has been determined to be about 30 μm, thecontainer should be scaled accordingly. The device further comprises apolarizable electrode in the container adjacent the target molecules.The electrode becomes polarized due to charge perturbation in thereaction mixture and target molecule. Further, an amplifying circuit formaintaining a set potential in the electrode and generating a signal inresponse to said charge perturbation; and a detector for detecting thesignal, thereby indicating reaction of the selected reactant with thetarget molecule are provided. A voltage clamp amplifier is exemplifiedas providing sensitivity to very small charge differences. The devicepreferably comprises an amplifying circuit with a differential feedbackamplifier having one input at a fixed voltage outside the container andanother input attached to the electrode.

The amplifying circuit preferably has its negative input connected tothe electrode, and the electrode is preferably gold, copper or silver.

The device may further comprise a self-assembled monolayer on thesurface of the electrode, for providing insulation, which willfacilitate polarization of the electrode. The self-assembled monolayeris preferably linked to the target molecule, so that the chargeperturbations on the target molecule occur within a zone that isdetectible by the electrode.

As stated, in the preferred device, the electrode is insulated, and thetarget molecule is a polynucleic acid (e.g. DNA or RNA) linked to thepolarizable electrode. The device may be fabricated as an array of alarge number of reaction electrodes, each with a different targetmolecule (e.g. DNA template). This device comprises an addressable arrayof electrodes and fluid circuits addressing each electrode individually.Also, the device may comprise an addressable array of electrodes andfluid circuits addressing the array collectively. That is, the fluidcircuits do not have to be designed to deliver separate reagents to eachelectrode. The same mixture (e.g. dNTPs) can be delivered to eachelectrode with a target molecule attached. A semiconductor substrate isprovided with integrated circuits for connections to electrodes and forthe amplifying circuits and channels for the fluid circuits.Microfluidics may be used, and the channels for delivering reactants maycomprise a polymer, e.g. poly(dimethylsiloxane) (PDMS) and/orPoly(silarylene siloxane) (PSS).

The present invention further includes methods for fabricating such adevice, and methods for using such a device.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is schematic diagram of a charge perturbation detection system;

FIG. 2 is a schematic showing the device of FIG. 1 as it operates in aDNA polymerization reaction;

FIG. 3 is a current trace of addition of a complementary nucleotide(3A); noncomplementary nucleotide (3B); and multiple nucleotides (3C);

FIG. 4 is a schematic drawing showing fabrication and surface chemistryof a preferred embodiment of the present invention;

FIG. 5 is a scheme of kinetic mechanism of nucleotide incorporation byDNA polymerase (5A); a graph of the relative induced charge (Qind/Q)function for the electrode surface geometry (x=0.3 mm, y=0.3 mm) andsource charge distance h (5B);

FIG. 6 is a diagram of a multiple electrode system useful for detectingpathogens;

FIG. 7 is a schematic diagram of a CMOS chip implementation of thepresent system;

FIG. 8 is a diagrammatic side section view of an integrated chargeperturbation device without a complete microfluidic channel (8A) andwith a microfluidic channel (8B); and

FIG. 9 is a diagram of a suitable low noise amplifier for use with anintegrated charge perturbation device.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT Introduction

The present device allows for direct detection of electric chargeperturbations of the local environment by an electrode system. When usedwith immobilized DNA, it can uniquely identify a DNA sequence. Thepolymerization process generates local perturbations of charge in thesolution near the electrode surface and induces a charge in thepolarazible electrode. This event is detected as a transient current bya voltage clamp amplifier. Detection of single nucleotides in a sequencecan be determined by dispensing individual dNTPs to the electrodesolution and detecting the charge perturbations. Alternatively, multiplebases can be determined at the same time using a mix of all dNTPs withsubsequent analysis of the resulting signal. The initial enzymeattachment to the DNA molecule can be detected prior to polymerizationwith a surface electrode capacitance measurement using the samevoltage-clamp amplifier. This technique thus enables also the detectionof enzyme-DNA interactions at the electrode surface.

Based on current understanding of polarizable interfaces, and withoutwishing to be bound by any particular scientific theory, it is believedthat the electrode is polarized, i.e. has a potential difference betweenthe exposed surface and the electrode, so that the negative chargeappearing in excess at the gold electrode surface is inducing the chargeof equal amount, but opposite in sign, on the other side of theinterphase without crossing the interfacial region. A negative inducedcharge on the side of the electrode, which is not facing the solutionwould be detected and amplified by a voltage clamp amplifier. Accordingto principles of electricity and magnetism, a charged object located inthe vicinity of a conducting surface will cause electrons to move in thesurface material even though there is no physical contact. This isillustrated in a simple electroscope, where a charged item held nearfoils causes them to separate. In this case, the charged item would bethe target molecule (DNA) and the electrode would become charged and fedinto the voltage clamp amplifier.

It is therefore possible to detect DNA sequences of target molecules bythe present Charge Perturbation Detection system by measuringcharacteristic transient current events by a voltage clamp amplifier.Signal or pattern recognition analysis of these events can be used todetermine the sequence of target molecules or identifies the targetmolecule itself.

Four different principal methods used by this technique are:

-   (1) Detection of the transient current events of the DNA    polymerization by adding individual dNTPs.-   (2) Detection of the transient current events of the DNA run-off    polymerization by adding a mix of all dNTPs to electrode solution.-   (3) Detection of the enzyme interactions (e.g. attachment) with DNA    molecule by measurement of capacitance changes of the electrode    interface.-   (4) Detection of a multiple DNA molecules and their mutations by    multi-electrode Charge Perturbation System.

A further aspect of this invention is the use of electrode coating withvarious layers to enable detection of DNA sequence by ChargePerturbation Detection system. The coating also enables and createsconditions for tethering a probe/template molecule to the electrodesurface.

The use of a template is designed to allow specificity of the reaction,as according to well-known rules of base pairing (Watson Crick pairing),A binds to T or U and G binds to C. The rapid and sensitive detection ofnucleic acids in a sample can provide a point-of-care diagnostic devicethat can test a patient's sample for the presence of specific knownpathogens and determine the proper course of treatment. This can also beuseful for current research applications that use DNA microarrays forreadout. Instead of labeling test samples with fluorescent molecules andhybridizing the DNA to microarrays, the samples can be hybridized toelectrode surface and detected with this system. This would result infaster and more accurate quantitative data with multiple sequencedetection in one test.

Charge Perturbation Detection of DNA Sequence Polymerization Concept:Introduction and Overview

The device described in detail below demonstrate that the electricalsignal created by polymerization of nucleic acids, and other biochemicalreactions is detectable and can be easily acquired using commercialelectronic components. Without extensive optimization, the sensitivityof the prototype was superior to that of commonly used optical detectionsystems such as real-time PCR systems.

CPD (Charge Perturbation Detection) technology does not require modifiednucleotides or marker molecules to detect polymerization signals.Through application of sensors that are implemented in CMOS technology,individual polymerization steps can be detected in real time and easilyprocessed. Not only does CPD avoid costly labels and specialized readoutequipment, but as noted above, preliminary results suggest that CPD isalso more sensitive. Because target DNA can be detected atconcentrations below 10 fmol, the present system does not needamplification of target DNA. Experimental systems have been fabricatedthat can detect 1-10 femtomoles (1−10×10⁻¹⁵ moles) of DNA, an order ofmagnitude higher detection sensitivity than commercial pyrosequencing orsingle-base extension technologies.

The underlying principle of CPD is the detection of the transientcurrent that is generated by the immobilization of charged moleculesonto a measurement electrode. In the case of DNA detection, transientcurrent is generated by proton removal as a complementarydeoxynucleotide-triphosphate (dNTP), incorporated into a nascent strandon the measurement electrode (upon which the target is immobilized witha probe that acts as a primer), and accompanied by proton removal duringbiomolecular reaction. The current signal is detected by the measurementelectrode. A characteristic current waveform is generated that can beinterpreted in terms of the polymerization rate of the participatingdNTP molecules. The resulting electrical signal is transient and is notpresent once the polymerization reaction is complete (i.e., over time,the rate of change in net charge becomes zero).

CPD differs from traditional electrochemical measurement techniqueswhere the applied potential difference speeds up the rate of oxidationor reduction in order to sense the rate of diffusion of a reactant inthe vicinity of the electrode surface. Our sensor applies a constantzero potential to the measurement electrode through a feedback loop tocompensate for accumulated charge at the electrode surface and toprevent diffusion of equilibrating counter-ions. The CPD detectionmethod actively senses transient current that results from changes inthe electrical charge during chemical reactions in the solution. Thechemical reactions that are detected are near, but not at theelectrode/solution interface, and no charge flows through the interface.Using the example of the incorporation of dNTPs, the signal issubsequently converted into a current signal by a high-gainvoltage-clamp amplifier. Upon sequential additions and washouts ofindividual dNTPs, nucleotide sequence peaks will appear as is shown in asequence poly-gram (FIG. 3). This simple procedure will identify anunknown sequence in the immobilized template strand.

The detection mechanism of the CPD devices are based on the chargeconservation principle, where the increase of the total negative chargeon the DNA molecules is exactly compensated by an increase of the totalpositive charge in the solution resulting from an increase of protonconcentration. Each of these electrical charges induces a surfacecharge, opposite in sign, on the coated, electrically isolated buthighly polarizable, gold electrode. The magnitude of any induced chargeis a function of the electrode surface geometry and the distance betweenthe electrode and the inducing charge. For electrodes used in theseexperiments, the magnitude of the induced charge is effectively constantfor separation distances in the range of ˜1 nm up to ˜30 μm (thedetection zone), and steeply decreases for distances >30 μm (FIG. 5B).The charges on a DNA molecule attached to the electrode are locallyfixed in close proximity to the electrode surface (<100 nm). The protonsreleased from DNA are free to rapidly diffuse in the solution far enoughto produce a change of the net charge in the detection zone. This eventinduces a charge sensed by the polarizable electrode. Since theelectrode is held at a constant potential, the charge induced by anindividual molecule results in a small pulse of current in theelectrode. The sum of these current pulses from all DNA moleculesattached to the electrode surface produces a large transient currentdetected by the voltage-clamp amplifier. Ideally, the measured currentis equal to the time rate of change in net charge within the detectionzone during the reaction, expressed by the equation I(t)=dQ(t)/dt (whereQ is charge, I is current and t is time). To evaluate actual efficiencyof signal transduction, surface DNA density was measured bypolymerization of radiolabeled dCTP. This showed that approximately 1femtomole of DNA was immobilized on the electrode surface (0.0009 cm²),which corresponds to ˜6.7×10¹¹ DNA molecules per cm² of the electrode.This result correlated well with the calculation of DNA surface densityof ˜6.0×10¹¹ molecules per cm² based on the size of the electronic CPDsignal of polymerization, indicating that the efficiency of signaltransduction of the CPD electrode is very high.

As discussed below in connection with FIG. 2, an incoming dNTP molecule,complexed with one Mg²⁺ ion (Ref. 15), increases the negative charge by2e⁻. Incorporation of the catalytic Mg²⁺ ion (Ref. 16) decreases thenegative charge by 2e⁻. Incorporation of a complementary nucleotide thenincreases a negative charge by 1e⁻ on the new backbone phosphate group,produced by removal of a proton from the 3′-OH group of the DNA primerduring the catalytic step of the reaction (Ref. 17), followed by rapiddiffusion of the proton into the surrounding solution. The change in theinduced charge can be detected by the electrode as a transient currentmeasured by a voltage-clamp amplifier. The diffusion distance oflow-molecular-weight compounds (Mg²⁺, MgdNTP²⁻, MgPPi²⁻) in solutionduring the time course of the experiment (1 s) is approximately an orderof magnitude slower than proton diffusion. For this reason the chargechanges induced by most of the reaction steps (binding of the dNTPmolecule, complexed with one Mg²⁺ ion, incorporation of the catalyticMg²⁺ ion, dissociation of the catalytic Mg²⁺ ion and of the leavingMg²⁺-bound pyrophosphate) do not produce a measurable electroderesponse. On the same basis, the Brownian motion of ions in the solutionas well as conformational changes of the immobilized enzyme and DNAmolecules do not produce changes in the induced charge.

In one embodiment, a prefabricated electrode matrix was used for DNAimmobilization. The electrically active 1×2 cm Au chips (shown in FIGS.1 and 4) were manufactured using a semiconductor-processing technique ona 4-inch wafer at the Stanford Nanofabrication Facility (SNF). The chipsconsisted of four pairs of rectangular gold electrodes that were 0.09mm² with 0.5 mm center-to-center spacing. For measurements of electricalactivity of the CPD electrode an Axopatch 200B voltage-clamp amplifier(Axon Instruments, Union City, Calif.) was used. The Axopatch amplifierwas used in the whole-cell voltage-clamp mode with the holding potentialat 0 mV. The high-impedance coating of the CPD measurement electrodeprevents the occurrence of faradaic current that could otherwise causeinterference or the deterioration of the sensor and analytes. Areference Ag/AgCl electrode of the voltage-clamp amplifier was immerseddirectly in the bathing solution during measurements.

For the initial experiments, single-stranded DNA molecules (76 bases)with different sequences were chemically synthesized with a thiolmodification on the 5′-terminus and HPLC purified by MWG Biotech (Highpoints, NC, USA). The DNA sequences were designed to self-prime with a19 base-pair self-complementary sequence at the 5′-end of the DNA.Approximately 40 bases of the DNA sequence are single stranded andextendable by DNA polymerase.

Self-priming, single-stranded DNA molecules were immobilized on thesurface of a gold electrode through a thiol-reactive self-assembledmonolayer (SAM). The electrode was equilibrated with 10 units of theKlenow (exo) fragment (KF) of DNA polymerase. FIG. 3A shows the signalresulting from the addition of a solution containing a single dNTP (1 mMconcentration in final solution volume) complementary to the nucleotidein the template sequence (top black trace).

The dNTPs used in the present system do not require any type oflabeling, nor is any reporter molecule used in the present system. Thisprevents problems associated with different incorporation parameters forlabeled dNTPs. Readouts of sequences of various lengths are thuspossible.

With no measurable delay, current rises to a peak of ˜400 picoamps (pA)within ˜50 milliseconds (ms), rapidly decreases to ˜50 pA, and thenshows a further, slower transient increase to ˜150 pA within 300 ms. Thecurrent transient is almost completed at 1 second (<5% of peak current).The integral of the measured current is 87 picocoulombs (pA·s),corresponding to nucleotide incorporation to ˜6.0×10¹¹ DNA molecules percm² of the electrode. In contrast, if a solution containing anon-complementary dNTP was added, no current transient was observed(FIG. 3B). No signal was produced when the complementary dNTP was addedin the absence of DNA polymerase, in the absence of DNA, or if DNA wasnot immobilized on the electrode surface. The lack of detectable signalin the control experiments demonstrates the clear dependence of thecurrent transient on the complementarity of the actual nucleotide, andon the simultaneous presence of KF (Klenow fragment) and immobilizedDNA. The current waveform observed can therefore be attributed to thesignal resulting from the incorporation of the nucleotide into theprimer strand.

Fabrication of Gold Chips

Electrically active 1×2 cm Au chips were manufactured using asemiconductor processing technique on a 4-inch wafer at the StanfordNanofabrication Facility (SNF) (http://snf.stanford.edu). The pairs ofelectrodes on the Au chips were arranged so that each electrode pair wasorthogonal to the adjacent pair, forming a generally square pattern ofelectrodes. Eight parallel leads were formed, extending from eachelectrode to the edge of the chip. The process requires only a singlemask, designed on an industry-standard CAD program and produced on apiece of Mylar thin-film. A 500 μm-thick quartz layer is used as thesubstrate. The process flow is as follows: A very thin layer of chromiumis first deposited to improve the adhesion between gold and quartz.Next, a 1000 {acute over (Å)} thick gold layer is deposited to definethe pattern for both the electrodes and the connecting pads. The minimumfeature size of this chip is 200 μm. To prevent contamination afterprocessing, a 7-μm-thick photoresist is used as a protection layer.After dicing, the photoresist is washed off with acetone andisopropanol. The chips consisted of four pairs of rectangular goldelectrodes that were 0.09 mm² with 0.5 mm center-to-center spacing.Other areas on the chip were used for the connection pads to externaldevices.

Surface Modification

All reagents used for surface modification were of reagent grade andused as received from Aldrich unless otherwise stated. The patternedquartz chips were cleaned in an RCA cleaning solution (H₂O: NH₄OH: 30%H₂O₂, 5:1:1, v/v/v) for 15 min at 70° C., immersed in a water bath for10 min and dried in a stream of argon. The quartz surface was coatedwith a hydrophobic octadecyltriethoxysilane (Gelest, Morrisville, Pa.)in an anhydrous toluene solution containing 1% (v) silane and 2% (v)hexanoic acid for 24 hr at room temperature. Silanized chips were washedtwice with toluene and once with ethanol for 5 min each, and dried in astream of argon. The silanization step was performed to make the quartzsurface hydrophobic and thereby avoid cross contamination between goldelectrodes in close proximity to each other during spotting.

The gold electrodes were coated with a long-chain thiol to form adensely packed monolayer (a self-assembled monolayer [SAM]), whichdisplaces any physisorbed silanes (Refs. 25, 26), i.e. silanesphysically absorbed on the surface. The silane-coated chips wereimmediately immersed in a 1 mM solution of mercaptoundecanol (MUD) inethanol for at least 16 hr. The gold substrates were removed from thethiol solution, washed with ethanol, and dried under an argon stream.The hydroxyl-terminated monolayer was transformed into a thiol-reactivemoiety by exposure to a 2.3 mM solution of N-(p-maleimidophenyl)isocyanates (PMPI, Pierce, Rockford Ill.) in anhydrous toluene at 40° C.for 2 hr under an argon atmosphere. (Refs. 27,28)

Suitable long chain thiols for linking to the electrode surface and thetest molecule (DNA strand) are generally between about 3 to 24 carbonlong. Generally, the electrode has bound to it an SAM linked to theelectrode at a thiol group at one end; a long chain alkyl group, and alinker group for attachment to a nucleic acid (DNA) at the other end. Asuitable linker group is PMPI.

Other suitable SAMs are described in U.S. Pat. No. 6,652,398, herebyincorporated by reference for purposes of describing SAMs containingphytanylthiol. As described in U.S. Pat. No. 6,048,623 to Everhart, etal., issued Apr. 11, 2000, entitled “Method of contact printing on goldcoated films,” incorporated by reference for further description ofsuitable SAMs, the self-assembling monolayer may have the followinggeneral formula:

X—R—Y where X reacts with the metal (gold) electrode, R is a linear orbranched alkyl or polymer backbone; and Y reacts with a nucleic acid,e.g. a thiol functionalized DNA. The DNA linkage may be through aphosphate to a sulfur group on a PMPI linker. The PMPI linker isattached to the SAM and to the DNA. PMPI(N-[p-Maleimidophenyl]isocyanate) is useful because both hydroxyl andsulfhydryl reactivity can be found in this cross-linker. Maleimidereacts with —SH groups at pH 6.5-7.5, forming stable thioether linkages.Isocyanate reacts with —OH groups to form a carbamate link at pH 8.5. Itis an excellent tool for conjugating —OH group-containing compounds.

X is designed to be reactive with metal or metal oxide. For example, Xmay be asymmetrical or symmetrical disulfide (—R′SSR, —RSSR), sulfide(—R′SR, —RSR), diselenide (—R′Se—SeR), selenide (—R′SeR, —RSeR), thiol(—SH), NH2, nitrile (—CN), isonitrile, nitro (—NO.sub.2), selenol(—SeH), trivalent phosphorous compounds, isothiocyanate, xanthate,thiocarbamate, phosphine, thioacid or dithioacid, carboxylic acids,hydroxylic acids, and hydroxamic acids.

Maleimide-modified gold electrodes were washed with anhydrous tolueneand dried in a stream of argon. The various surface modification stepswere followed by X-ray photoelectron spectroscopy (data not shown) andthe presence of the expected elements and peak shifts confirmed theproper transformation of both surface components.

An SAM may also be formed of polymers that bind to gold and or otherelectrode materials. For example, Major, et al. “Strategies for CovalentMultilayer Growth,” Chem. Mater. 2002, 14, 2574-2581 describes astrategy for the covalent assembly of polymer multilayers at interfaces,where growth is accomplished one layer at a time. The individual layerconstituents are maleimidevinyl ether alternating copolymers with sidegroups that possess reactive functionalities. The identity of thepolymer layer side groups determines the chemistry employed ininterlayer linkage formation. It is reported there that one may carryout the selective creation of amide, ester, ether, urea, and urethaneinterlayer linkages.

The SAM serves to provide insulation for the electrode and to preventnon-specific adhesion of DNA to the electrode. The SAM facilitates theorientation of the DNA in a parallel linkage to the substrate.

Immobilization of DNA

The thiolated oligonucleotides were diluted to a final concentration of10 μM in 0.1M phosphate buffer, pH 7.4, with 10 μM dithiothreitol (DTT)and incubated at least 1 hour at room temperature. Immobilization of thereduced thiolated oligonucleotides onto the electrodes was performedmanually by deposition of 0.2 μl reduced oligonucleotides followed byovernight incubation at room temperature in a humidified chamber.

Design, Synthesis and Purification of Oligonucleotides

Single-stranded DNA molecules (76 bases) with different sequences werechemically synthesized with a thiol modification on the 5′-terminus andHPLC purified by MWG Biotech (High points, NC, USA). The DNA sequenceswere designed to self-prime with a 19 base-pair self-complementarysequence at the 5′-end of the DNA. Approximately 40 bases of the DNAsequence are single stranded and extendable by DNA polymerase. Theoligonucleotide used in the experiment (FIG. 1) was:

(SEQ ID No: 1) 5′-Thiol/TTTTTTTTTTTTTTTTTTTTGCTGGAATTCGTCAGTGACGCCGTCGTTTTACAACGGAACGG CAGCAAAATGTTGC.

Prototype Sensor System

In a typical charge-based sensor system, a prefabricated electrodematrix is used for DNA immobilization. The electrodes were fabricated asdescribed above. The CPD electrode chip that was used in theseexperiments is further described in the section “fabrication of goldchips.”. The electrode surface was submersed in a standard DNApolymerization buffer (5 mM Tris-HCl, pH 8.3; 25 mM KCl and 1.25 mMMgCl₂) with DNA polymerase (10U, Klenow fragment). Polymerization wasinitiated by adding a 20 aliquot containing 20 mM of dNTP substrate (togive 1 mM final concentration of dNTP in 400 buffer).

Electrical Measurement Method

For measurements of electrical activity of the CPD electrode was used anAxopatch 200B voltage-clamp amplifier (Axon Instruments, Union City,Calif.). The Axopatch amplifier was used in the whole-cell voltage clampmode with the holding potential at 0 mV. The high-impedance coating ofthe CPD measurement electrode prevents the occurrence of faradaiccurrent that could otherwise cause interference or the deterioration ofthe sensor and analytes. A reference Ag/AgCl electrode of thevoltage-clamp amplifier was immersed directly in the bathing solutionduring measurements.

Other Amplifier Designs

A voltage clamp amplifier may described generically as a differentialfeedback amplifier having one input at a fixed voltage outside thecontainer and another input attached to the electrode. Manygeneral-purpose op amp chips have two or four separate operationalamplifiers in one package, with common power supply connections. Inpractice the ideal amplifier criteria requirements are met onlyapproximately, but as will be shown, close enough for most purposes.Practically, an op amp will have a gain of 10,000 or more, an inputimpedance of megohms, and a 3 dB bandwidth of several tens of hertz ormore. If an amplifier has a 3 dB bandwidth of 40 Hz and a gain of100,000 times, this is a gain bandwidth product of 4 million hertz, or 4MHz (40×100,000). It is advantageous in many feedback applications tohave the gain falling at 6 dB per octave or 20 dB per decade atfrequencies beyond the corner frequency (that frequency at which theamplifier gain has fallen 3 dB or 70.7 percent of its DC value). Sincethe op amp is used in mainly in feedback circuits having much lowerclosed loop gain, these performance figures are good enough in manycases. In fact, even a single high gain (100×) common emitter transistoramplifier stage can be treated as an op amp if feedback is employed,with surprisingly little error. In many cases a single transistor willwork almost as well as a more expensive op amp device. One example is asimple audio amplifier stage from which a moderate gain (5-20×) isrequired.

Radiolabeling

Radioactive labeling and phosphor imaging techniques were used asprocess controls to quantify the oligonucleotide attachment andsubsequent hybridization reactions.29 [a-32P] dCTP (Pharmacia) was usedfor 3′ labeling of the attached self-primed probes via single-baseextension. Specific activities of the radiolabeled oligos weredetermined by liquid scintillation counting using an LS 7500 liquidscintillation system (Beckman Inc., Columbia, Md.). Standard curves weremade from a serial dilution of known amounts of the 32P-labelednucleotides used in the experiments. The data presented here representthe averages of, minimally, three replicate points.

Example 1 (FIG. 1): Schematic Diagram of Charge Perturbation DetectionSystem

FIG. 1 is a schematic diagram of a device according to the presentinvention, with attached enzyme and substrate, showing the generation ofions and the detection circuitry. The device is formed on a substrate10, by the attachment of a metal electrode 12, preferably of anoxidation resistant, high potential metal such as gold. A coating orinsulation layer 14, may also be applied to the electrode 12. A templateDNA strand 16 is attached to the electrode structure. A primer DNAstrand is hybridized to the template DNA strand 16. Deoxynucleotidetriphosphate (dNTP) and Mg²⁺ ions are added. Incorporation of thecomplementary dNTP releases inorganic pyro phosphate (PPi) and H⁺ ions(protons). A DNA polymerase complex is illustrated at 17 and ionmovement is shown relative to the enzymatic reactions, i.e. dNTP, MG²⁺in and PPi and H⁺ out.

As described below, the charge perturbation in the vicinity of theelectrode is detected by a voltage clamp amplifier 18 having a negativelead 20 connected to the electrode 12 and a positive lead connected to acommand voltage source 22. Feedback resistor 24 serves to eliminatecurrent output whenever the command voltage is equal to the referencevoltage at the measurement electrode. When a small charge perturbationoccurs at the electrode 12, an amplified, transient signal 26 isproduced. A reference electrode, which does not have a template DNAstrand attached to the electrode structure can be also provided (but isnot required). The reference electrode is connected to a similar op amp18 a with feedback resistor 24 a to provide a reference signal forcomparison against the transient signals from the reaction electrode.

Prior to the start of the polymerization reactions, a DNA template isimmobilized on a coated electrode surface. DNA polymerization requiresthe hybridization of a sequencing primer to this strand, so that DNApolymerization with attached enzyme can proceed with dNTP dispensationsto the electrode solution and the sequence of the template strand can bedetected. The system can also be designed to detect the hybridization ofa template strand that is specific for a sample to be analyzed, in whichcase the sequence of the template strand is already known and thereaction detects the presence of a complementary primer strand in thesample to be analyzed.

Example 2 (FIG. 2): Template Orientation and Operation

DNA polymerase catalyzes the incorporation of complementary nucleotides,resulting in perturbations of local electrical charge. Perturbations arebased on interactions of system components and release of products ofpolymerization. These local perturbations of charge in solution induce acharge in coated polarazible gold Charge Perturbation Detection system(CPD) electrode. This event is detected as a transient current byvoltage clamp amplifier, which compensates induced charge in the goldelectrode. The high-impedance coating prevents faradaic current thatcould deteriorate the sensor and analytes. For measurements ofelectrical activity of CPD electrode was used Axopatch 200B voltageclamp amplifier, but in principle any voltage-clamp amplifier can beused. Voltage command was kept at steady OmV holding potential, butbecause no faradaic current is present the voltage command can be kepton different holding potentials and can be used for various purposes inthe final system. For example, a specific voltage can repel variouspolymerization products or ions and “clean” the surface of the electrodefor next polymerization steps.

FIG. 2 shows base pairing that has already occurred at positions 202 and204. Template strand base 206 is about to be paired with a base 206 a,which includes triphosphate. Two of the phosphates are stabilized with amagnesium ion, as is known to be required by DNA polymerase. Arrow 208shows the dNTP with base 206 a prior to incorporation; arrow 210represents incorporation of base 206 a. Base 206 a is then linked to thegrowing (primer) strand through the phosphate backbone. At that point,there is a new charge incorporated into the primer strand phosphatebackbone, and, as shown by arrow 212, there is liberated H⁺, PPi(inorganic pyro phosphate) and catalytic Mg²⁺. The H⁺ diffuses away andan electron is repelled from the electrode 12 interfacing solution,causing a positive induced charge in the electrode, as shown by arrow220, and a negative current pulse 222 that is detected by the sensitiveamplifier 18. The sensitivity is provided by a voltage clamp circuit andamplified to an output signal 26.

Capacitance Measurement

Measurement of the electrode capacitance is used to detect attachment ofan enzyme to the DNA located at the surface of the electrode. This stepwill easily eliminate problematic electrodes and speed up measurement ofelectrodes with correct enzyme/DNA interaction. Capacitance increase ofelectrode interface is probably caused by the change of DNA dielectriclayer due to the enzyme attachment to the DNA. If there is a problem ofenzyme/DNA attachment or in the electrode surface chemistry fabrication(e.g. thickness of coated layers) then no significant increase incapacitance is observed even all components are present in the solution.Increase of capacitance in current version of electrodes (size, layers,enzyme, buffer concentration, DNA density) is about 5-6 fold and candepend on mentioned parameters. Without wishing to be bound by onetheory as to the cause of this effect, it is thought that targetmolecules (DNA strands) with attached enzyme “stand up,” i.e. arealigned away from the electrode, allowing solution to be closer to theSAM, effectively making the dielectric thinner.

Measurement of capacitance is based on applying small (e.g. ±2 mV)triangle-wave voltage command to the electrode and detecting responsesquare-wave current recalculated to capacitance by equation:C=I/(dV/dt). Measurement of electrode capacitance enables also detectionof general interactions of enzymes with the target molecule (DNA)located at the surface of the electrode.

Example 3 (FIG. 3): Analysis of Signals

If the addition of individual dNTPs is performed sequentially, thennucleotide sequence peaks will appear as shown in a sequence poly-gram(FIG. 3). The individual detected signal is shown in FIG. 3A and FIG.3C. If the addition of dNTPs is performed with a mix of all nucleotidesthen a run-off polymerization signal is detected (FIG. 3C). This signalcan be then analyzed and multiple nucleotide bases can be determined bythe shape of the curve, which will vary in response to the particularnucleotide incorporated.

In FIG. 3 A, the addition of a dGTP binding to a C nucleotide in thetemplate is shown. The spikes at the beginning of the signal are causedby dispensation disturbance and initial diffusion of dGTP solution andare easily filtered out. With no measurable delay after dispensation,current rises to a peak of ˜400 picoamps (pA) within ˜50 milliseconds(ms), rapidly decreases to ˜50 pA, and then shows a further, slowertransient increase to ˜150 pA within 300 ms. The trace in FIG. 3B showsthat no signal is generated in the absence of a reaction near theelectrode interface.

Example 4 (FIG. 4): Electrode Fabrication and Chemistry

Fabrication of a “CPD chip” (Charge perturbation detection chip) on aglass wafer is illustrated in FIG. 4, in which the steps in a surfacechemistry process to produce coated electrodes with immobilized DNA areoutlined. The size of the illustrated electrode is 0.3×0.3 mm² and thedistance between two electrodes in each electrode pair is 0.2 mm. Mostpart of the chip excluding the electrodes and a part of the paths iscovered by a dielectric layer (silicon oxide) to avoid direct contactbetween the paths and the solution. The number of electrodes is notlimited by the design and it will vary in individual applications,number of sequences to be detected and/or micro fluidics to be used. Theelectrodes described in FIG. 4 are manufactured usingsemiconductor-processing technique on a 4-inch wafer (FIG. 4) at theStanford Nanofabrication Facility (SNF). The process requires only asingle mask that is designed using an industry-standard CAD program andproduced on a piece of mylar thin-film. A 500 μm thick quartz layer isused as the substrate. The process flow works as follows: A very thinlayer of chromium is first deposited to improve the adhesion betweengold and quartz. Next, a 1000 {acute over (Å)} thick gold layer isdeposited to define the pattern for both the electrodes and theconnecting pads. The minimum feature size of this chip is 40 μm. Toprevent contamination after processing, a 7 um thick photoresist is usedas a passivation layer.

In step 410, the gold electrode is cleaned with a standard cleaningprocess:

1) UV Ozone cleaner, from Jelight Company Inc. (www.jelight.com)

2) Plasma oven from 4th state inc (www.4thstate.com)

3) RCA cleaning, containing water, ammonium hydroxide, hydrogen peroxide(5:1:1).

In step 420, a silane layer is applied to areas surrounding theelectrode. In step 430, a self-assembled monolayer is attached to theelectrode areas where the DNA template will be immobilized. Next in situformation is carried out in step 440. This creates a reactive group forthe attachment of a sulfur-containing DNA strand in step 460. That is,after attaching the monolayer 442 linked to the gold, a layer ofPM-1-N-(pMaleimidophenyl) isocynate 444 is used to attach covalentlythiol modified DNA 462 to the PMPI.

CPD electrodes coated as described above may be placed into a holder inorder to contain the chip and provide fluid paths and electricalconnections. Electrode output is connected to the Axopatch 200Bamplifier or equivalent voltage-clamp device (Axon Instruments, FosterCity, Calif.). The sample solution (e.g. 400), consisting of buffer (5mM Tris-HCL, pH 8.3 (at 25° C.); 25 mM KCl and 1.5 mM MgCl₂) isdispensed into an electrode chamber.

Example 5 Operation of Illustrated Embodiment

Self-priming, single-stranded DNA molecules were immobilized on thesurface of a gold electrode through a thiol-reactive self-assembledmonolayer (SAM). The electrode was equilibrated with 10 units of theKlenow (exo-) fragment (KF) of DNA polymerase.

Initial capacitance of the electrode was measured by triangle-wavevoltage, as described above under “Capacitance Measurement.” A sawtoothwaveform is applied to the command voltage. Then polymerase enzyme (e.g.2 μl) is introduced to the solution. After several minutes, electrodecapacitance is measured. Correct attachment of enzyme is confirmed byincrease of electrode capacitance. The amount of increase depends fromthe size of electrode and DNA density and should be several times higherthan initial capacitance without enzyme (e.g. 5×). Finally individualdNTPs, or a mixture of dNTPs, are rapidly dispensed to the solution.Speed of diffusion and dispensation should be in the range of severalmilliseconds in order to avoid disturbance of following polymerizationreaction. Piezo-spray dispensation triggered by computer (piezo, RJ315with Sone-Tek generator) was used for dispensation, but some otherfast-perfusion system with for example air-pressure spritzer can be usedas replacement (e.g.

Spritzer-8 8-Channel Micro Injector—Spritzer from BioSciense Tools,Inc.). Concentration of dNTPs can be scaled according to signalsensitivity. Signal to noise ratio is based on size of the electrode,DNA density, solution concentration, etc.). We used for example 5 mM-20mM sol of dNTP.

Transient ionic current is recorded using the Axopatch 200B amplifier involtage-clamp mode with signal filtering at 5-10 kHz bandwidth. Thesignal is further digitized by an Axon Digidata 1320A digitizer withsampling frequencies from 10 kHz to 500 kHz. The data is recorded usingClampex 8 (Axon Instruments), and the same software is used for basicsignal analysis.

The signal resulting from the addition of a solution containing a singledNTP (1 mM concentration in final solution volume) complementary to thenucleotide in the template sequence rises to a peak of about 400picoamps (pA) within ˜50 milliseconds (ms), rapidly decreases to ˜50 pA,and then shows a further, slower transient increase to ˜150 pA within300 ms. The current transient is almost completed at 1 second (<5% ofpeak current). The integral of the measured current is 87 picocoulombs(pA·s), corresponding to nucleotide incorporation to ˜6.0×10¹¹ DNAmolecules per cm² of the electrode. In contrast, if a solutioncontaining a noncomplementary dNTP was added, no current transient wasobserved. No signal was produced when the complementary dNTP was addedin the absence of DNA polymerase, in the absence of DNA, or if DNA wasnot immobilized on the electrode surface (data not shown; see priorityprovisional). The lack of detectable signal in the control experimentsdemonstrates the clear dependence of the current transient on thecomplementarity of the actual nucleotide, and on the simultaneouspresence of Klenow fragment (KF) and immobilized DNA. The currentwaveform observed can therefore be attributed to the signal resultingfrom the incorporation of the nucleotide into the primer strand. The DNApolymerase-catalyzed elongation of the synthesized strand proceeds bythe SN2 (bimolecular nucleophilic substitution) mechanism that has beenextensively studied. (Refs.9, 10, 11). Upon the incorporation of eachnucleotide, the total negative electrical charge on the DNA moleculeundergoes a net increase of 1e⁻, produced by the removal of a protonfrom the 3′-OH group of the DNA primer during the catalytic step of thereaction. (Ref. 12).

Example 6 (FIG. 5): Electrochemistry Reactions

Due to the principle of charge conservation, the increase of the totalnegative charge on the DNA molecules is exactly compensated by anincrease of the total positive charge in the solution resulting from anincrease of proton concentration. Each of these electrical chargesinduces a surface charge, opposite in sign, (Refs.3, 4) on the coated,electrically isolated but highly polarizable, gold electrode. Themagnitude of any induced charge is a function of the electrode surfacegeometry and the distance between the electrode and the inducing charge.For electrodes used in these experiments, the magnitude of the inducedcharge is effectively constant for separation distances in the range of˜1 nm up to ˜30 μm (the detection zone), and steeply decreases fordistances >30 μm (FIG. 5B). (Refs.3, 4)

FIG. 5B shows the relative induced charge (Q_(ind)/Q) for the electrodesurface geometry (x=0.3 mm, y=0.3 mm) as a function of the distance hbetween the electrode and the charge Q in the solution (Ref. 4). In therange of ˜1 nm-˜30 μm, the electrode response does not depend on thedistance between the electrode and the ion. Therefore, the changes inthe immobilized charge can produce a change of the induced charge on theelectrode only if the released countercharge diffuses to a distanceof >30 μm from the electrode surface, for the present size electrode.

The charges on the DNA molecule attached to the electrode are locallyfixed in close proximity to the electrode surface (<100 nm), while theprotons released from DNA are free to diffuse in the solution. For theduration of the experiment (˜1 s), the diffusion distance of protonsis >136 μm (the diffusion coefficient of proton D_(H) ⁺ in water is9.3×10⁻⁵ cm²/s) (Ref. 13). Lateral proton diffusion might besignificantly faster due to specific surface hydration of the electrodeand the surrounding silane layer. (Refs. 14,15) The protons are thusable to diffuse far enough to produce a change of the net charge in thedetection zone due to the immobilized negative charge on the DNAbackbone. This event induces a charge sensed by the polarizableelectrode. Since the electrode is held at a constant potential, thecharge induced by an individual molecule results in a small pulse ofcurrent in the electrode. The sum of these current pulses from all DNAmolecules attached to the electrode surface produces a large transientcurrent detected by the voltage-clamp amplifier. Ideally, the measuredcurrent is equal to the time rate of change in net charge within thedetection zone during the reaction, expressed by the equationI(t)=dQ(t)/dt (where Q is charge, I is current and t is time). Toevaluate actual efficiency of signal transduction we measured surfaceDNA density by polymerization of radiolabeled dCTP, as discussed above,indicating that DNA polymerization with as little as 1 fmole of DNA on asingle electrode could be detected. Less than 1 fmole of DNA could beused as a template if a smaller electrode were used. To gain furtherinsight into the mechanisms underlying CPD during a single step of theKF-catalyzed reaction, the kinetics of the reaction were studied. (Refs.11, 16, 17, 18, 19, 20, 21)

The kinetic scheme based on these studies is shown in FIG. 5A, whichuses the following abbreviations: E: enzyme, Ea: active enzyme, Mg:magnesium, Dn: DNA with enzyme at n-th nucleotide, dNTP: deoxynucleoside5′-triphosphate, PPi: pyrophosphate, kn: kinetic rate. The kineticmechanism of nucleotide incorporation by DNA polymerase is showninitially to the left of arrows (1) (step 1). The enzyme forms initiallya binary complex with the DNA primer-template (E·Dn). Upon addition ofdNTP to the solution, the initial binding of an incoming complementarydNTP to polymerase produces a ternary substrate complex (step 1). Afterthis step, the enzyme undergoes a subdomain motion to form a so-calledclosed or active state (step 2), which is followed by binding of thecatalytic Mg²⁺ ion (step 3) and by the chemical reaction of dNTPincorporation onto the DNA primer strand by the formation of aphosphodiester bond (step 4). Unbinding of the catalytic Mg²⁺ ion (step5) is followed by a second subdomain motion of the polymerase/productcomplex, resulting in an open state (step 6) followed by the release ofpyrophosphate PPi (step 7) and subsequent DNA translocation.

Thus, the incoming dNTP molecule, complexed with one Mg²⁺ ion, (Ref. 22)increases the negative charge by 2e⁻. Incorporation of the catalyticMg²⁺ ion (Ref 23) decreases the negative charge by 2e⁻ (1) Incorporationof nucleotide (2) then increases a negative charge by 1e⁻ on the newbackbone phosphate group (3), produced by removal of a proton from the3′-OH group of the DNA primer during the catalytic step of the reaction(Ref. 12), followed by rapid diffusion of the proton into thesurrounding solution (4). The change in the induced charge (5) can bedetected by the electrode as a transient current (6) measured by avoltage-clamp amplifier (7). The diffusion distance oflow-molecular-weight compounds (Mg²⁺, MgdNTP²⁻, MgPPi²⁻) in solutionduring the time course of the experiment (1 s) is approximately an orderof magnitude slower than proton diffusion (Ref. 24). For this reason thecharge changes induced by most of the reaction steps (binding of thedNTP molecule, complexed with one Mg²⁺ ion, incorporation of thecatalytic Mg²⁺ ion, dissociation of the catalytic Mg²⁺ ion and of theleaving Mg²⁺-bound pyrophosphate) do not produce a measurable electroderesponse. On the same basis, the Brownian motion of ions in the solutionas well as conformational changes of the immobilized enzyme and DNAmolecules does not produce changes in the induced charge.

Sample kinetic rates, using the notation of FIG. 5A, for a trace such asshown in FIG. 4A are as follows:

k _(on)=1.2×10⁷ M⁻¹s⁻¹,

k _(off)=0.06 s⁻¹,k₁=1.25×10⁷ M⁻¹s⁻¹,

k ⁻¹=250 s⁻¹,

k ₂=50 s⁻¹,

k ⁻²=3 s⁻¹,

k ₃=9.5×10⁵ M⁻¹s⁻¹,

k ⁻³=100 s⁻¹,

k ₄=150 s⁻¹,

k ⁻⁴=40 s⁻¹,

k ₅=100 s⁻¹,

k ⁻⁵=9.5×10⁵ M⁻¹s⁻¹,

k ₆=4 s⁻¹,

k ⁻⁶=4 s⁻¹,

k ₇=60 s⁻¹,

k ⁻⁷=1.45×10⁴ M⁻¹ s⁻¹.

Based on this kinetic model, a simulation was performed to confirmwhether this underlying mechanism accounts for the signal dynamicsobserved. The initial enzyme-DNA binding step was assumed to be inequilibrium during the simulation, since the enzyme was incubated formore than 2 minutes with the DNA attached to the electrode in theexperiments. The published rate constants of individual reaction stepswere used as the starting point of simulations, with the exception ofstep 3 and 5, for which we found no published KF rate constants. Weapproximated these steps with Kd˜100 μM based on the published rateconstants for DNA polymerase β (Ref. 19). We achieved a best fit betweenmodel and experimental signals by adjustment of selected kinetic rates(in particular, k₆, k⁻⁶, k₇, k⁻⁷ were changed somewhat from thepreviously published values to account for different experimentalconditions) (Refs. 20, 9). Using Arndt's model (Ref. 21), anapproximation for the rate of binding of Mg²⁺ and with the adjusted rateconstants, we have recreated the key features of the signal dynamics,mainly the timing of the peaks of polymerization, which closely match inboth the simulation and the experiment.

In summary, the charge perturbation detection concept is based on thecharge conservation principle and the induced surface charge of thepolarizable electrode. As a result, the immobilized negative chargeaccumulated on the DNA backbone can be detected as soon as the positiveproton leaves the detection zone, which occurs in a relatively shortperiod. In principle, DNA synthesis confined to the detection zone couldbe detected without DNA immobilization, since the protons diffuse fasterthan the DNA molecules. The described label-free detection method can beeasily applied to a multiple-electrode system and can provide rapid andsensitive detection of biological pathogens, genetic mutations oridentification of unknown DNA sequences with a very small amount ofsample. In addition, it also enables general measurements of enzymesundergoing similar catalytic reactions. The CPD electrode thuspotentially represents a very robust and effective biosensor for manymolecular and diagnostic applications.

Example 7 (FIGS. 6 and 7): Arrangement of CPD Devices into DetectionArrays

Multi-electrode DNA detection can target a specific list of DNAsequences, such as for example DNA mutations of a pathogen. The conceptis shown in FIG. 6. In the first panel, a series of leads 602individually connect each of six electrodes to a detection circuit,comprising a sensitive voltage clamp circuit as described above. Eachelectrode has immobilized on it a single DNA strand (primer) 603 from aspecific strain of human papilloma virus (HPV), e.g. HPV6 (electrode 1),HPV16 (electrode 2), HPV11 (electrode 3), HPV18 (electrode 4), HPV31(electrode 5), and HPV45 (electrode 6). Appropriate DNA sequences to beimmobilized as primers may be obtained from GenBank, e.g. Locus 573503,Locus HPV16R, Locus HPV 11R, etc., available through NCBI Entrez system,online. Locus identifications are as of the date of last modificationlisted in that entry as of the filing date of the present application.HPV is a double stranded DNA virus, whose entire genome has beensequenced. Over 80 strains of HPV are known.

In the second panel of FIG. 6, a sample DNA 604 with an exposed singlestrand DNA region from a sample containing HPV is added to the array.The sample should be at least partially denatured to allow DNAhybridization. In the third step, the sample has hybridized to the HPVimmobilized DNA 603 and DNA polymerase enzyme 606 and dNTPs are added.In the fourth step, only the hybridized target is polymerizing andgenerating a signal 608.

Appropriate primers may be made by standard PCR protocol. Due to thesensitive nature of the present method, less than about 10 femtomol ofprimer is needed per electrode. However, more primers may be added to asingle electrode to facilitate assembly.

In addition, multiple primers for different regions of DNA for a singletarget (e.g. different HPV6 genes) may be added to a single electrode toimprove sensitivity. Multiple-electrode chips can find specificmutations in one-step run-off polymerization or by sequentialdispensations and detection of individual nucleotides.

Referring now to FIG. 7, the present device may be integrated on a CMOSchip. A CMOS (complementary metal oxide semiconductor) semiconductoruses both NMOS (negative polarity) and PMOS (positive polarity)circuits. The chip 710 contains within it the differential amplifiers712, 714 that are connected to each electrode. An insulating layer 716contains connecting lines and vias 718 for connecting each electrode toan amplifier. A pad 720 is provided for a solder ball 722 to connect toan aluminum connector 724, which passes through a quartz or glass wafer726. An electrode 728 of gold connects to the connector 724. Theelectrode has a chemical functionality 730 for immobilizing the testmolecule (e.g. DNA) 732 to the electrode. A reference electrode 734 doesnot contain test molecules.

The electrode array is electrically connected via solder balls to theCMOS semiconductor chip. The electrode array is constructed on a quartz,glass or silicon wafer. The sensors, amplifiers and signal routing willbe embedded in the CMOS chip and the output signal will be accessiblevia a pad and/or solder ball.

Example 8 (FIGS. 8-9): Integrated Designs

An exemplary integrated detection device is illustrated in FIGS. 8 and9. This design implements further a design according to FIG. 7. Thisdevice can be used for parallel detection of DNA targets in an arrayformat. The system presented here can detect low (femtomolar) amounts ofDNA molecules, obviating the need for DNA amplification. It can beapplied to the analysis of gene expression, short DNA sequencing, SNPdetection, and pathogen detection.

When larger numbers of functional areas and a small device size isdesired, one may use a modified microarray contact printer equipped witha digital camera and controlling software for precise solution deliveryto the individual electrodes. Alignment of the contact pins with theelectrodes will require some calibration and testing to ensure properalignment. Testing may be done by optical image processing.

Another critical issue is the surface chemistry and attachment of theDNA. During the fabrication process, an oxide layer can form thatprevents efficient attachment of the DNA oligonucleotides. Usingelectrical measurements of surface capacitance, as described above, onemay determine the efficiency of attachment of the DNA-enzyme complex.Alternative methods include semi-manual dispensing of reagents by Piezospray and XPS analysis to ensure that the chip is functional.

A microfluidic device according to the present invention may employ anautomated method comprised of two components to accomplish a fastperfusion of buffer for equilibration and washing, and a more accuratedispensing of reagents.

The fast perfusion (˜100 ms) of polymerization buffer in a desktop unitmay be done with a Valvelink 8.2 controller and pinch valves (AutomatedScientific Inc., CA), controlled by computer with an NationalInstruments (NI) 6225 board. The accurate injection (˜10 ms) of 1-20 ofdNTPs to the CPD electrode is performed with calibrated glassmicropipettes connected to an eight-channel Micro Injector-Spritzer(BioScience Tools, CA) and positioned closely to the CPD electrode witha multi-electrode holder (BioScience Tools) and MM-33 micromanipulator.

For final assembly of the automated desktop system, the newlyfunctionalized and tested chip will be connected to the reagentdispensing system. The chip will be placed in a prefabricatedchip-holder that is connected to the multiple voltage-clamp amplifiers(Dagan, Inc.).

For sensor data acquisition, commercially available data acquisitionboards and software packages from National Instruments are used. TheirLabView software package can be used for any necessary processing on thePC. The NI 6225 board will be used. The NI 6225 can sample 16single-ended channels at a rate of 15.6 kHz per channel. The boardallows to be multiplexed so that up to 80 channels can be sampled at arate of 3.1 kHz.

The DNA probe is first printed on the electrode chip and inserted intothe device. The solution containing the target is added and allowed tohybridize to the probe (not shown). Then solution buffer (50 mM NaCl,1.5 mM Mg), and DNA polymerase is introduced into the electrode chamberat the beginning of each measurement cycle. The individual dNTPs aredispensed sequentially (one dNTP per cycle) to the solution by using thefast perfusion/dispensation system. Transient ionic current signal ineach cycle is detected using the voltage-clamp amplifier and digitizedby computer acquisition board to produce the sequence polygram.

Microfluidic technology to deliver analytes and other small volumes ofliquid can be adapted to a larger (e.g. an 8×8) biosensor array. Thetechnology has sub-micron alignment precision of the microfluidicchannels on top of the biosensor array. Furthermore, the microfluidictechnology is compatible with materials that cannot withstand etching,elevated temperatures, or high electric fields.

Suitable microfluidic technology is formed of two layers and containsmicrofluidic channels interconnected with 8 parallel channels. Eachchannel runs over 8 sensors. The electrical wires and contacts arearrayed orthogonally to the microfluidic channels. The microfluidicchannels may be designed to deliver the same mixture of reactants, orsequential reactants, to each electrode, or, with more fabrication, toprovide individual channels for delivering different reactants todifferent electrodes.

Each of the 64 sensors can be measured independently. A thick top layer,fabricated by soft lithography from polydimethylsiloxane (PDMS) andbacked by a transparent glass wafer, defines the larger microfluidicstructures, such as inlet and outlet channels. Small microfluidicstructures that would be difficult to create and align if they wereincluded in the PDMS layer are instead defined in a thin layer of SiO₂,which has been deposited by ion beam deposition (IBD) and patterned by alift-off process on top of the existing biosensor array. The combinationof IBD and lift-off is compatible with the delicate biosensors andprovides photolithographic alignment precision and feature definition ofthe channels in the SiO₂ layer. Moreover, the IBD-deposited SiO₂ layerprovides spontaneous adhesion and sealing to the top PDMS layer and willaccommodate a significant amount of substrate topography. Large channeloverlaps between the two layers make the alignment of the top PDMS layerto the substrate error-tolerant. Finite element simulations are used tooptimize the two-layer geometry for high analytic sensitivity and lowflow impedance.

Microfluidic channels are formed as channels in parallel in onedirection, e.g. from top to bottom of integrated device. Then, theelectrodes to the sensor points would go from left to right. The sensorlocations are at the intersection of the microfluidic channels andelectrodes, forming an 8×8 array.

Fluidic channels down to approximately 2 um width have been created withgood edge definition in a 250 nanometer thick SiO₂ layer, spanning 5 μmacross a sensor area between intersections with larger channels in thePDMS layer.

Referring now to FIG. 8A, a cross section of a biochip is shown, afterthe SiO₂ channels are built on the chip and prior to DNAfunctionalization of the electrodes. A silicon substrate 810 hasembedded therein an aluminum lead, which is part of a network of leadsconnecting individually each gold electrode in an array 814. The Sisubstrate 810 has an SiO₂ insulating layer 816, except for the exposed,functionalized portion 818. The exposed portion has an SAM 820. DNAmolecules 822 are applied by a robotic dispenser or spotter 824.

DNA is applied at the wafer level by selecting a particular DNA solutionfor each electrode. Details of the biofunctionalization chemistry aredescribed above.

For a compact charge perturbation device, the electrode size ispreferably about 0.3 mm by 0.3 mm, which can be accommodated by existingassembly devices. The 0.3 mm electrode size can be scaled down tomicrometer scale electrode with the use of semiconductor processingtechnology and microfluidics.

FIG. 8B shows the cross section of the biochip in 8A after sealing theDNA probes with PDMS channels 826 and illustrates how the electrodes fora CPD are integrated with microfluidics. The fabrication of preciselyaligned microfluidic channels with micron-scale features on theelectrodes has been simplified by combining the sealing andconformational benefits of PDMS soft lithography with the precision anddurability of photolithographically patterned SiO₂. Thebiosensor-compatible microfabrication process provides reliable sealingon various substrate topographies. The tolerance in aligning theelastomeric PDMS layer to the substrate makes automation feasible, whichshould facilitate the development of inexpensive diagnostic biochips.

The microfluidic channel is sealed with a quartz sheet 828 on top of thePDMS layer 826. The channel is thus defined by a bottom surfacecontaining DNA molecules adjacent an electrode; sidewalls comprising adeposited polymeric layer; and a top sheet confining liquid containing asample to be analyzed and analytical reagents (e.g. polymerase, dNTPs,etc.) Two channels are illustrated in FIG. 8B.

An integrated circuit version of the CPD sensor using CMOS technologyhas several advantages over a version built from off-the-shelfcomponents. First, it has increased detection sensitivity and increasedsignal-to-noise ratio because it can be optimized for low-noiseoperation. Second, it can be cheaper in volume production than a versionbuilt with off-the-shelf components. Third, it can be smaller than aversion built with off-the-shelf components.

The signal-detection circuitry consists of two stages. The first stageis a voltage-clamp amplifier 910, which amplifies and converts currentto voltage. The second stage 912 is used for output sampling in order toreduce noise. It performs a correlated double sampling of the firststage output. The correlated double sampling technique commonly used indigital cameras and in image sensor technology, cancels the lowfrequency 1/f noise, which is introduced by the first stage amplifier.Low frequency 1/f noise is inversely proportional to frequency so thatthe lower the frequency, the stronger the noise component. In commercialCMOS processes, the 1/f noise becomes equal to flat white noise in thekHz range. By canceling the 1/f noise with the correlateddouble-sampling technique, sensor sensitivity can be significantlyimproved by an order of magnitude.

The charge integrating amplifier 914 contains a ground and an input 916,along with a feed back resistor R, as is standard in voltage clampamplifiers. The circuit further comprises a capacitor C for chargeintegration and a feed back switch (D_(i). The output of the chargeamplifier is input to either a positive or negative input to anamplifier 918. Switches Φ₂ and Φ₃ connect the charge amplifier to thedifferential amplifier. The switches Φ₁₋₃ operate sequentially, i.e. oneis on, then two is on, then three is on. This provides an over frequencyto the output to be amplified, and, as explained above, lowers noiselevels.

The amplifiers are designed using transistors in standard topologies andare optimized for low noise. Optimization of the charge-integratingamplifier circuit is possible after the impedance of the ssDNA-modifiedelectrode is measured. Optimization will primarily consist of selectionof the feedback impedance in the charge amplifier to maximize signalgain.

Integrated circuit design and verification software are indispensableparts of the design process. The basic software tools are schematiccapture, parasitic extraction, circuit simulation, mask layout, anddesign verification. Industry-standard software is available for suchuses, e.g. as provided by Cadence Design Systems for schematic capture,mask design, parasitic extraction, and design verification. The softwareHSpice may be used for circuit simulation.

Differential Amplifiers

The present devices utilize amplifiers that can detect small differencesof charge at their inputs. The classic design for such an amplifier isthe “voltage clamp” design illustrated here. An op-amp with no feedback(i.e. no resistors 24, 24 a in FIG. 1) is already a differentialamplifier, amplifying the voltage difference between the two inputs.However, its gain cannot be controlled, and it is generally too high tobe of any practical use. The present application of negative feedback toop-amps has resulting in the practical loss of one of the inputs (e.g.the ground), the resulting amplifier is only good for amplifying asingle voltage signal input. However, using known amplifier designprinciples, one can construct an op-amp circuit maintaining both voltageinputs, yet with a controlled gain set by external resistors. Otheramplifiers and circuits may be constructed to allow for the control ofthe gain and sensitivity of the amplifier circuit. Buffer amplifiers aretypically added to improve performance.

Example 9 Detection of Other Reactions

The sensitivity of the present CPD system enables the detection of otherreactions besides DNA polymerization. Polymerization of other nucleicacids may also be detected in a sequence specific manner. The system maybe used to detect binding of RNA to DNA, or double stranded RNAmolecules found RNAi and ribozyme constructs. Furthermore, otherbiological molecular reactions generate more or less charged products,which can be detected using the described devices.

For example, the cleavage or unfolding of a polypeptide to exposecharged arginine, lysine, histidine, aspartate, or glutamate residuesmay be detected. A protein may be immobilized on an electrode and asmall amount of a protease and a potential protease inhibitor added, sothat the effectiveness of a panel of protease inhibitors may bedetermined in numerous parallel reactions.

Molecules that act on phosphate groups may also be monitored by thepresent CPD system. This includes phosphotases and kinases. Measurablekinases include protein kinases, such as protein kinase C andserine/threonin kinsases, as well as pyruvate kinases, glycerol kinasesand the like.

Other ATP dependent reactions measurable in the present system includethat activity of DNA gyrase (which is modified by certain antibiotics)and the synthesis of glutathione (gamma-glu-cys-gly; GSH). GSH isusually present at high concentrations in most living cells, being themajor reservoir of non-protein reduced sulfur. Because of its uniqueredox and nucleophilic properties, GSH serves in bio-reductive reactionsas an important line of defense against reactive oxygen species,xenobiotics and heavy metals. GSH is synthesized from its constituentamino acids by two ATP-dependent reactions catalyzed bygamma-glutamylcysteine synthetase and glutathione synthetase.

Other enzymes, such as desulfurases (e.g. cysteine desulfurase),generate protons in solution and their activity will be detectible andmeasurable by the present devices. Sulfonucleotide reductases are adiverse family of enzymes that catalyze the first committed step ofreductive sulfur assimilation. In this reaction, activated sulfate inthe context of adenosine-phosphosulfate (APS) or-phosphoadenosine-phosphosulfate (PAPS) is converted to sulfite withreducing equivalents from thioredoxin. The sulfite generated in thisreaction is utilized in bacteria and plants for the eventual productionof essential biomolecules such as cysteine and coenzyme A. Humans do notpossess a homologous metabolic pathway, and thus, these enzymesrepresent attractive targets for therapeutic intervention. Inhibitors ofsuch enzymes could be added to wells containing one or more variants ofsuch enzymes and their sulfate substrates and cofactors. Further detailsare given in Carroll K S, Gao H, Chen H, Stout C D, Leary J A, et al.(2005) “A Conserved Mechanism for Sulfonucleotide Reduction,” PLoS Biol3(8): e250.

Other applications for measurement of reactions with the present methodsand apparatus include the measurement of the release or formation ofphosphate drugs, such as pyrophosphate analogs, e.g. 4-amino-1hydroxybutylidene bisphosphonic acid monosodium salt (Fosamax®). Cellmembranes containing ion channels may be attached to the electrodes, andthe effect of test substances on ion permeability measured.

The present DNA sequencing reactions could also be repeated using theoriginal templates, if the hybridized strand is denatured and washedaway, leaving the target DNA. This would allow re-use of a device.

The present examples, methods, procedures, specific compounds andmolecules are meant to exemplify and illustrate the invention and shouldin no way be seen as limiting the scope of the invention, which isdefined by the literal and equivalent scope of the appended claims. Anypatents or publications mentioned in this specification are indicativeof levels of those skilled in the art to which the patent pertains andare intended to convey details of the invention which may not beexplicitly set out but would be understood by workers in the field. Suchpatents or publications are hereby incorporated by reference to the sameextent as if each was specifically and individually incorporated byreference and for the purpose of describing and enabling the method ormaterial referred to.

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What is claimed is:
 1. A sensor comprising: an electrode; an insulating layer on a surface of the electrode; at least one template nucleic acid coupled to the insulating layer; and circuitry to apply a first voltage to the electrode, and to detect a current pulse through the electrode in response to incorporation of known nucleotides into the at least one template nucleic acid.
 2. The sensor of claim 1, wherein the known nucleotides are in a solution contacting the insulating layer.
 3. The sensor of claim 1, wherein the circuitry includes a feedback loop to apply the first voltage to the electrode.
 4. The sensor of claim 1, wherein the current pulse is induced in response to charged particles liberated during the incorporation.
 5. The sensor of claim 1, wherein the current pulse is induced in response to a change in surface charge of the insulating layer due to the incorporation.
 6. The sensor of claim 1, wherein the at least one template nucleic acid is attached to the insulating layer.
 7. The sensor of claim 1, wherein the known nucleotides are label-free.
 8. The sensor of claim 1, wherein the current pulse indicates at least a portion of a sequence of at least a portion of the template nucleic acid.
 9. The sensor of claim 1, further comprising a fluidic structure to provide the known nucleotides.
 10. The sensor of claim 1, further comprising a second electrode to provide a reference signal, and wherein the circuitry compares the detected current pulse to the reference signal to detect occurrence of the incorporation.
 11. A method for detecting a chemical reaction, the method comprising: providing a template nucleic acid into a container communicating with an electrode via an insulating layer; applying a first voltage to the electrode; introducing known nucleotides into the container; and detecting a current pulse through the electrode indicating incorporation of the known nucleotides into the template nucleic acid.
 12. The method of claim 11, further comprising analyzing the current pulse to determine at least a portion of a sequence of at least a portion of the template nucleic acid.
 13. The method of claim 11, wherein the known nucleotides are label-free.
 14. The method of claim 11, wherein the template nucleic acid is attached to the insulating layer.
 15. The method of claim 11, wherein the current pulse is induced in response to a change in charged particles within the container due to the incorporation.
 16. The method of claim 11, wherein the current pulse is due to a change in surface charge of the insulating layer due to the incorporation.
 17. The method of claim 11, further comprising: receiving a reference signal from a second electrode; and comparing the detected current pulse to the reference signal to detect occurrence of the incorporation.
 18. The method of claim 17, wherein the second electrode is not coupled to a second template nucleic acid.
 19. A sensor comprising: an electrode; an insulating layer on a surface of the electrode; at least one template nucleic acid coupled to the insulating layer; a circuit to apply a first voltage to the electrode; and a detector to detect a current pulse through the electrode in response to incorporation of known nucleotides into the at least one template nucleic acid.
 20. The sensor of claim 19, wherein the known nucleotides are in a solution contacting the insulating layer.
 21. The sensor of claim 19, wherein the circuitry includes a feedback loop to apply the first voltage to the electrode.
 22. The sensor of claim 19, wherein the current pulse is induced in response to charged particles liberated during the incorporation. 